Protein Delivery Applications
4.2 Polymers for the Design of Hydrogels
4.2.4 Thermogelling Polymer Hydrogels
Thermosensitive polymers belong to the class of stimuli-responsive materials, also known as “smart,” “intelligent” or “environmentally sensitive”. Thermoresponsive poly-mers exhibit differences in solubility in aqueous medium in response to temperature changes. The temperature at which they undergo this transition is commonly referred to as the lower critical solution temperature (LCST). Below the LCST, the polymers are soluble in aqueous medium, whereas above the LCST, they are insoluble, due to hydro-phobic interactions between the polymer chains. When thermosensitive and perma-nently hydrophilic polymers are combined (in the form of block or graft copolymers), they are able to self-assemble above their LCST, forming a hydrogel structure. Polymers with LCST between room and body temperature can be used for the preparation of injectable hydrogels. In Figure 4.4 the self-assembly mechanism of a thermosensitive triblock copolymer is depicted.
Figure 4.4 Self-assembly of a thermosensitive linear triblock copolymer composed of two outer thermosensitive blocks and an inner hydrophilic block yielding the formation of a hydrogel above the LCST of the thermosensitive block.
In-Situ Gelling Thermosensitive Hydrogels for Protein 101 Temperature-responsive polymers are characterized by a critical gelation temper-ature in aqueous solutions where self-assembly of the polymer chains occurs owing to hydrophobic interactions, thus phase separation is observed. Polymers can display Lower Critical Solution Temperature (LCST) or Upper Critical Solution Temperature (UCST) when the polymer solution is phase separated above or below a specific tem-perature, respectively. Thermosensitive behavior is generally viewed as a phenomenon governed by the balance of hydrophilic and hydrophobic moieties on the polymer chain [40,41]. Most of the polymers studied for biomedical applications exhibit LCST behavior, with a few exceptions, for example, natural polymers like gelatin and polysac-charides like agarose or cellulose derivatives [42,43]. Only a few natural polymers dis-play LCST behavior in the range between room and body temperature. Some cellulose derivatives (methyl and hydroxypropyl methylcellulose) at low concentrations (1–10 wt%) are liquid at low temperature, but jellify upon heating, however their gelation temperature is far above body temperature, representing a limitation of this material as in-situ gelling systems [44]. Chemical and/or physical modification can be adopted to lower the gelation temperature, for example, by addition of NaCl or decreasing the hydroxypropyl molar substitution of hydroxypropyl methylcellulose [44,45]. However, no studies on protein release with these systems have been published to date. Chitosan has been reported by Chenite et al. to form a gel close to body temperature and at physiological pH when combined with glycerol phosphate disodium [46]. Bhattarai et al. developed an injectable, thermoreversible gel based on chitosan-PEG copolymer (chitosan-g-PEG), that utilized intermolecular chitosan chain interactions for gelation.
This hydrogel was used as a depot system for sustained protein release [47]. This type of thermosensitive gelation has also been observed in cellulose derivatives grafted with hydrophilic moieties [2]. Synthetic polymers offer many more opportunities as com-pared to natural polymers for the design of injectable hydrogels. The most frequently studied synthetic thermsosensitive polymer for biomedical and pharmaceutical appli-cations is poly(N-isopropylacrylamide) (PNIPAM), because its LCST in water is 32°C, thus suitable for in-situ gelling (Figure 4.5) [41]. The incorporation of hydrophilic monomers in pNIPAM increases the LCST, whereas more hydrophobic units decreases it [48]. Similar behavior was observed by Vermonden et al., who reported a decrease in LCST of poly(hydroxylpropyl methacrylamide lactate) (pHPMAm-lac) upon intro-duction of hydrophobic methacrylate moieties in the polymer lactate side chains [23].
Similarly, the gelation behavior of poly(d,l-lactide-co-glycolide)–poly(ethylenglycol)–
poly(d,l-lactide-co-glycolide) (PLGA-PEG-PLGA) was influenced by the hydrophobic-ity of end caps (hydroxy, acetyl, propionyl, and butanoyl groups); an increase in the hydrophobicity of the copolymer lowered the transition temperature [49]. The same finding was obtained with cholesterol end-capped star PEG-PLLA copolymers [50].
Physically crosslinked PNIPAM-based hydrogels were described for the first time by Han et al. [51], who synthesized poly(N-isopropylacrylamide-co-acrylic acid) (p(NIPAM-co-AA)) to prepare thermosensitive matrices that were used in follow-up studies for biomedical purposes, particularly as synthetic matrices in refillable bio-artificial pancreas. Encapsulated Langeran islets showed good viability, and the cell-laden artificial matrices showed insulin release [52,53]. Similarly and more recently, p(NIPAM) networks were crosslinked using N, N′-methylenebisacrylamide (BIS) and used for bovine serum albumin (BSA) release studies in vitro. The release of the protein
102 Handbook of Polymers for Pharmaceutical Technologies
was not complete and a strong interaction between polymer and protein was proposed as the reason for the non-retrieved protein [54]. Kim et al. described the use of pH/
thermosensitive polymeric beads based on terpolymers of NIPAM, butyl methacrylate (BMA) and acrylic acid (AA) (pH-sensitive) to modulate release of insulin. A high loading efficiency was accomplished (90–95%), and while no release of insulin was observed at pH 2.0 and 37 °C, the drug was released at physiological pH [55]. The release rate and mechanism depended on the molecular weight (MW) of the polymer:
low MW terpolymers eroded very quickly and released insulin within 2 hours by an erosion mediated mechanism, while in the case of high MW polymer, which had a better stability, the gels showed a release of insulin for 8 hours that was governed by swelling/diffusion [55]. As observed for many other thermally assembled polymers, the stability of such hydrogels is rather poor and represents a major limitation in the use of these materials for pharmaceutical purposes. Therefore, in recent years, strat-egies to improve the stability of thermosensitive networks by chemical crosslinking methods, suitable for in-situ gelling, have been exploited. PNIPAM-based hydrogels self-assembling in a thermoreversible fashion and displaying improved hydrophilicity, thus an enhanced capability to retain water within the hydrogels matrix, were synthe-sized by grafting NIPAM to permanently hydrophilic polymers like poly(ethylenglycol) (PEG), for example, via Ce4+/OH redox initiated free radical polymerization [56,57]. A series of polymers with different architectures were synthesized (AB, BAB, A(B)4, and A(B)8 linear and star-shaped block copolymers with PEG as A block and PNIPAM as B block), and characterized for gelation mechanism and exploited for chondrocytes immobilization [58]. Several other copolymers of PNIPAM with poly(2-methacryloy-loxyethyl phosphorylcholine) (PMPC) were synthesized and characterized [59,60].
Other non-degradable thermosensitive polymers exhibiting hydrophilic-hydrophobic transition at temperatures close to body temperature are poly(vinyl ether)s (PVEs);
their derivatives and copolymers [61] are excellently reviewed elsewhere [62] and are beyond the scope of this chapter, as data on pharmaceutical and biomedical applica-tions to date are lacking.
A series of polymers, namely Pluronic® (BASF), based on poly(ethylene oxide)-b-poly(propylene oxide)-b-poly(ethylene oxide) triblock copolymers (PEO–PPO–PEO), with varying PEO/PPO molecular weights and contents, exhibit LCST behavior below body temperature [63] and have been extensively investigated for their physical-chem-ical and thermodynamic properties, and pharmaceutphysical-chem-ical applications [28,64–66].
Pluronics have been extensively used as in-situ forming drug delivery matrices and the possibility to prolong to some extent the drug pharmacokinetics by using Pluronic-based hydrogels was demonstrated. For example, similarly to NIPAM, monoamino-ter-minated Pluronic (mainly Poloxamer PF127) was coupled to poly(acrylic acid) (PAA) using dicyclohexyl carbodiimide (DCC) and graft copolymers of poly(acrylic acid)-g-Pluronic of different MW were synthesized via chain transfer reactions [67]. These graft copolymers gave improved gelation and mechanical properties as compared to the corresponding Pluronic, due to the presence of pH-sensitive moieties (PAA) that affect ionization and chain expansion of the polymer. Pluronic-based copolymers were widely studied for the delivery of protein and peptide therapeutics, like insulin, inter-leukin-2, urease, epidermal growth factor, endothelial cell growth factor, etc. Sustained release over several hours was observed with a possibility to tailor the release kinetics
In-Situ Gelling Thermosensitive Hydrogels for Protein 103 by polymer concentration or addition of excipients [68–72]. However, Pluronics, as well as pNIPAM, are not ideal biomaterials for in-vivo applications. Besides toxicity issues, observed with Pluronics in intraocular implantation [73], their main disad-vantage is their non-biodegradability that makes surgical intervention necessary to remove the delivery system from the body after the drug has been released. In addi-tion, weak mechanical strength and stability, as well as high permeability for entrapped compounds are further limitations associated with the use of these polymers. Some of the listed drawbacks were partially overcome; for example, Cohn et al. copolymerized PEG and PPO segments using two synthetic pathways: 1) chain extension of native Pluronics with hexamethylene diisocyanate (HDI) and 2) covalent binding of PEG and PPO chains using phosgene as the connecting molecule. The multiblock copolymers synthesized displayed remarkably improved mechanical properties as compared to Pluronic, moreover an extension of the drug release time (release of RG-13577 up to 40 days) as compared to self-assembled Pluronic hydrogels [74]. However, biodegradabil-ity issues still exist [75]. Many block copolymers of Pluronics with polyesters (PLA and PCL) were also reported [76,77]. The most advanced thermosensitive delivery systems for proteins rely on biodegradable polymers, which is very advantageous for in-vivo applications. Biodegradable and biocompatible PEG/polyester block copolymer hydro-gels, initiated by Kim and coworkers [78], were introduced in 1990s as a novel class of biodegradable thermosensitive matrices. ABA-type PEG–poly(l-lactide)–PEG triblock copolymers (PEG–PLLA–PEG) were first synthesized by ring-opening polymerization of l-lactide (LLA) using the monomethoxy PEG (MPEG) as macroinitiator; then the PEG–PLLA–PEG triblock copolymers were obtained by coupling MPEG–PLLA using hexamethylene diisocyanate (HMDI). These polymers exhibited UCST, therefore the drug-loaded hydrogels were prepared at 45 °C and then gelation was induced by low-ering the temperature below 37 °C. The release of fluorescein isothiocyanate (FITC) labeled dextran was studied and it was demonstrated that 12 days sustained release was achieved for 35 wt% hydrogels. Formulations of lower polymer content showed burst release that could be decreased by increasing the polymer concentration. Also, a series of star-shaped PLLA–PEG block copolymers were synthesized by coupling star PLLA with monocarboxy-MPEG using DCC coupling reaction [79]. The main disadvantages of this system are the long degradation time due to PLLA crystallinity and the need for high temperatures for the preparation of the hydrogels, as the polymer exhibits UCST behavior. Under these conditions the structure of labile protein molecules, along with their activity, might be affected.
The next generation of PEG/polyesters hydrogels were based on PEG-PLGA-PEG triblock copolymers [80]. These materials displayed both LCST and UCST behav-ior and were processable avoiding the use of high temperatures to dissolve the poly-mer. The hydrogel stability upon subcutaneous injection in vivo was demonstrated using rat models and one month stable matrices were obtained [81]. TGF-β1 was loaded into these hydrogels and used as a reservoir for controlled drug release aimed at wound healing purposes [82]. Significantly high levels of re-epithelialization, cell proliferation and collagen organization were observed. The sustained release of syn-thetic drugs like ketoprofen and spirolactone was also studied from PEG-PLGA-PEG hydrogels [20], as well as release of insulin, porcine growth hormone and glycosylated granulocyte colony-stimulating factor (in vitro and in vivo) [83,84]. Taken together,
104 Handbook of Polymers for Pharmaceutical Technologies
all these studies confirmed an improved stability and capability to release drugs over an extended period of time (weeks) as compared to Pluronic formulations. Several other thermosensitive copolymers of PEG with aliphatic polyesters were synthesized and applied for drug delivery. Some examples are AB, ABA, BAB copolymers of PEG with caprolactone (PCL) and δ-valerolactone (PVL) [52,85–87]. The vitro and in-vivo release of fluorescein isothiocyanate-labeled bovine serum albumin (BSA-FITC) was studied from PEG-PCL diblock copolymer gels and compared to Pluronic gels, demonstrating longer in-vivo stability of PEG-PCL hydrogels and enhanced capability to provide sustained protein release over 10 days, as compared to Pluronic gels, where destabilization and drug release within 3 days was observed [21,84]. Mikos et al. pro-posed the synthesis of PEG-based triblock copolymers consisting of poly(propylene fumarate) (PPF) as middle block [88]. Compared to other PEG copolymers, the use of PPF has the advantage of having unsaturated double bonds suitable for the stabilization of the hydrogels by chemical crosslinking of the hydrogels. Biodegradable multiblock amphiphilic and thermosensitive poly(ether ester urethane)s consisting of poly-[(R)-3-hydroxybutyratel (PHB), poly(ethylene glycol) (PEG), and poly(propylene glycol) (PPG) blocks were synthesized by Loh et al. Their aqueous solutions were found to undergo a reversible sol-gel transition by micellar packing upon temperature changes at very low copolymer concentrations (2–5 wt%) and the authors envisioned that these systems are suitable for protein delivery [89]. Recently, Pluronic analogs containing middle blocks of poly(hexamethylene adipate) (PHA), poly(ethylene adipate) (PEA), and poly(ethylene succinate) (PESc) instead of PPO were synthesized. Because of the hydrophobic nature of PHA and PEA, strong hydrophobic interactions and micelliza-tion occurred, leading to formamicelliza-tion of hydrogels only at high concentramicelliza-tions, while the more hydrophilic PESs showed gelation at low concentrations [90]. A general drawback of some of these polyester-based copolymers is their very long degradation time (i.e., PCL degrades in vivo in 2 to 4 years [91]) that leads to polymer accumula-tion in the body for long periods, limiting the use of these materials for chronic dis-eases, where controlled delivery systems need to be administered repeatedly. In 2004, Hennink et al. introduced a new class of thermosensitive and biodegradable polymers based on pHPMAm-lac that displays tunable LCST behavior from ~ 10 to 60°C by simply changing the length of the lactate side chains [17]. The polymer biodegradabil-ity is ensured by the presence of hydrolytically sensitive ester bonds in the lactate side chains. When the terminal lactate group is cleaved by hydrolysis, the resulting polymer becomes water soluble and can be eliminated by renal filtration, as long as its molecular weight is lower than the renal cutoff [92]. These thermosensitive polymers have been coupled to PEG by free radical polymerization using a PEG macroinitiator and yield a copolymer with ABA triblock architecture, consisting of inner PEG B-block flanked by outer p(HPMAm-lac) A-blocks. These polymers are suitable for the preparation of in-situ gelling systems [93], whose mechanical properties and degradation behavior were improved by combining thermal self-assembly with photopolymerization upon poly-mer derivatization with methacrylate moieties [23]. The chemically stabilized hydro-gels were suitable as controlled protein delivery, where model proteins were released according to diffusion-governed kinetics, easily tailorable from 1 week to 2 months by changing polymer molecular weight, concentration and degree of derivatization with methacrylate groups [12,15,24,94,95]. The potential of this thermosensitive hydrogel
In-Situ Gelling Thermosensitive Hydrogels for Protein 105 for tissue engineering was assessed by demonstrating good viability and differentiation of human Mesenchymal Stem Cells (hMSCs). Emerging thermosensitive hydrogels in the field of protein delivery are also biodegradable polyphosphazenes, consisting of a hydrophilic PEG block and hydrophobic amino acids or a peptide block (L-isoleucine ethyl ester (IleOEt), D,L-leucine ethyl ester (LeuOEt), L-valine ethyl ester (ValOEt), or di-, tri-, and oligo-peptides in the side groups [22,96,97]. Hydrogels were formed by intermolecular association of hydrophobic peptide chains and when PEG was coupled to di-, tri-, and oligo-peptides as side groups, hydrogels of higher mechanical strength were obtained as compared to PEG-IlaOEt polymer gels. Polymers containing depsi-peptide (GlyGlycOEt) showed faster hydrolytical degradation because of the genera-tion of carboxylic acid groups that made the polymers more hydrophilic, resulting in sustained release FITC dextran and human serum albumin in vitro for about 2 weeks.
The authors also studied strategies to decrease the burst release from polyphosphazene hydrogels by addition of chitosan that due to its positive charge retained negatively charged proteins like BSA, gelatin type B (GB20), and FITC-BSA within the hydrogel network. Application of these hydrogels as extracellular matrix for artificial pancreas was investigated [98,99].
Finally, polypeptides are important biodegradable and biocompatible polymers hav-ing a variety of conformations, such as α-helix, β-sheet, and random coil and they lend themselves to the synthesis of building blocks for the preparation of thermosensitive hydrogels with potential biomedical applications. Tirrell et al. introduced a polymer consisting of leucine zipper terminated protein flanking a central, flexible, water-soluble polyelectrolyte segment. Formation of coiled-coil aggregates of the terminal domains in near-neutral aqueous solutions triggers formation of a three-dimensional polymer net-work, with the polyelectrolyte segment retaining solvent and preventing precipitation of the chain [100]. Kopeček et al. reported a hybrid hydrogel system assembled from water-soluble synthetic polymers and a coiled-coil protein-folding motif. These hydro-gels underwent temperature-induced collapse owing to the cooperative conformational transition of the coiled-coil protein domain [101]. Shortly after, amphiphilic diblock copolypeptide that assembles into a gel both by supramolecular and thermal associa-tion were reported by Nowak et al. [102]. Jeong et al. developed L/DL poly(alanine) (PA) end-capped poly(propylene glycol)-poly(ethylene glycol)-poly(propylene glycol) (PLX) (PA-PLX-PA) polymers that in aqueous solutions underwent a sol-to-gel transi-tion at increasing temperature by increase in β-sheet content of PA and dehydratransi-tion of PLX. This system was stable in vivo for over 15 days [103]. Physically crosslinked poly(amino acid) hydrogels, formed by a sol-gel transition of amphiphilic poly(N-sub-stituted α/β-asparagine)s in an aqueous solution, were described by Takeuchi et al. [104], while recently Jeong et al. described amphiphilic polymers consisting of the hydrophilic poly(N-vinyl pyrrolidone) (PVP) block and a hydrophobic poly(alanine) (PA) block that formed micelles in water which aggregated as the temperature increased to yield gels.
They demonstrated the use of PVP as an alternative to PEG to design reverse thermo-gelling biomaterials [105]. Copolymer hydrogels that are pH and temperature sensitive were prepared by introducing pH-sensitive moieties in a temperature-sensitive polymer.
A pH/thermo-sensitive ABA copolymer was obtained by introducing carboxylic acid groups end groups into PLGA–PEG–PLGA triblock copolymers. Although the non-modified triblock copolymer did not exhibit gelation upon increase of temperature, the
106 Handbook of Polymers for Pharmaceutical Technologies
carboxyl-capped PLGA–PEG–PLGA led to four states (sol, gel, precipitate, and turbid sol) depending on pH and temperature [106]. Multiblock poly(ester amino urethane) s that are pH and temperature sensitive were synthesized by coupling poly(amino ure-thane) (PAU) through a condensation reaction to PCL-PEG-PCL triblock copolymers to yield multiblock copolymers (PCL-PEG-PCL-PAU)n. The incorporation of the ion-izable PAU segments in the macromolecule induced pH sensitivity due to the tertiary amine moieties. Thus, below pH 6.9 the polymer is in a sol state in aqueous solution up to 60 °C due to the electrostatic repulsion of the piperazine groups. In contrast, at physiological pH (pH 7.4) the solution displays a sol–gel transition upon increasing tem-perature to 37 °C. The formation of the free-standing gel depended on the formation of interconnected micelles. The formation of a gel was assessed in vivo [107] and injectable poly(amidoamine)-poly(ethylene glycol)-poly(amidoamine) triblock copolymer hydro-gels exhibiting pH and temperature sensitivity were designed for bioadhesive appli-cations. The dual responsiveness depended upon the poly(amidoamine) outer blocks which turned from a hydrophilic into hydrophobic state upon increasing pH and/or temperature. At low pH a sol was observed up to 60 °C, while above pH 7.0 the micelles bridged, leading to the formation of a gel. In-vivo experiments showed that upon sub-cutaneous injection of 12.5 wt% copolymer solution a white gel was obtained after one minute [108]. A thermosensitive triblock copolymer composed of poly(ε-caprolactone-co-lactic acid)-PEG-poly(ε-caprolactone-poly(ε-caprolactone-co-lactic acid) (PCLA-PEG-PCLA) was syn-thesized by ring-opening polymerization using ε-caprolacton (CL), lactide (LA) and PEG as macroinitiator. Separately, carboxylic acid terminated sulfamethazine oligomers (OSMs) were polymerized by chain transfer polymerization and coupled to terminal hydroxyl groups of the triblock copolymer, yielding a pentablock copolymer (OSMs-PCLA-PEG-PCLA-OSMs). The synthesized polymer solution showed a reversible sol-gel transition by a small pH change in the range of pH 7.4–8.0 and also by a temperature change in the region of body temperature, forming a gel at 37 °C, pH 7.4. The block copolymers OSM−PCLA−PEG−PCLA did not form a gel at pH 8.0 in the tested temper-ature range (from 4 to 60 °C) because the hydrophobic interaction between PCLA−OSM blocks is perturbed by the ionized sulfonamide group of the OSM block. As the pH is decreased, most of the OSM is deionized, restoring the hydrophobic interaction between PCLA−OSM blocks and forming a gel. By exploiting both pH and thermosensitive func-tionalities of the polymer, it was possible to broaden the gel window and obtain a sol between 10 and 70° C at pH 8.0. It is clear that this solution could be injected without concerns for premature gelation in the needle and, once in the body, the physiological pH triggered the gel formation. PCLA-PEG-PCLA gels showed a drop of pH (from 7.4 to 2.2), whereas the pentablock OSMs-PCLA-PEG-PCLA-OSMs showed the capability to buffer the pH, maintaining a value of 5.5. [109,110]
The same group synthesized diblock copolymer hydrogels based on a basic poly(β- aminoester) (PAE) coupled to MPEG. A gel-to-sol transition at pH > 6.0 was observed when the temperature was increased as a result of micelle packing [111]. Another dual-responsive polymer, PAE-PCL-PEG-PCL-PAE pentablock copolymer, used for the release of insulin, was prepared by Michael addition polymerization of 4,4-trimethylene dipiperidine (TMDP), PCL-PEG-PCL diacrylate, and butane-1,4-diol diacrylate (BDA).
Insulin, loaded into the hydrogels, formed complexes with the polymer, lowering its LCST and acting as physical crosslinks, which was confirmed by the longer stability
In-Situ Gelling Thermosensitive Hydrogels for Protein 107
of the protein-loaded hydrogels as compared with the placebo gels [112]. Most of the thermosensitive polymer hydrogels described in this section are overviewed in Table 1.1.